1. Field of the Invention
The disclosure in the present application is in the field of 3-D volumetric medical imaging, and, in particular, is directed to an improved device and method to determine tissue properties in the body of a subject.
2. Description of the Related Art
Many diagnostic 3-D imaging devices are used in modem medicine for many types of exams, which are performed by subjective viewing of the images on either film or more commonly on electronic displays. This subjective viewing uses assumed quantitative image pixels, which define boundaries of tissues, organs or foreign masses, and discrimination of tissue types based on grey scale or color images, representative of the underlying tissue properties. Identification of diagnostic details is fundamentally dependent upon the detection of image detail edges. Diagnostic interpretation or measurements from these images assume an appropriate relationship of the image to the tissue property. All of these imaging devices however have significant limitations for quantitative measurements because the relationship is not sufficiently accurate or defined. The embodiments disclosed in the present application are directed at improving these limitations while reducing costs, time and effort in the clinical setting.
Measurement, display, and analysis of tissue properties from medical images have many diagnostic benefits in the living subject. These include such measures as density, mass, volume, edges, etc.; image display reference values, translucency, texture; tissue water or hydrogen content, electron density, blood vessel iodine contrast media concentration and density or blood flow, tissue iron content, fat/muscle ratio, air/tissue ratio, and the like. Modern imaging devices enable the potential to identify and quantify, either automatically or manually, most organs and tissues of the body. The imaging devices to which the improvements disclosed herein are directed provide a set of volumetric images acquired from multiple projection angles, which allow reconstruction of images in various planes. Such devices include dual-energy and single-energy CT scanners, rotational C-arms, MRI devices, x-ray tomosynthesis and 3D DXA (dual-energy X-absorptometry) in which the x-ray source is moved to provide various projections, such as a Hologic 3-D DXA device in development. Several tissue measurements of interest include bone density, lung nodule density, cardiac, aortic and vascular calcifications, vascular soft plaque, fat measurements, muscle mass, lung volume and density (emphysema and the like), liver iron content, perfusion and blood flow, organ volume, density and mass, contrast angiography and the like. Tissue corrections for radiotherapy dose calculations are based on the electron density. Accurate and conveniently available measures of true tissue properties and their change with disease conditions or therapy allow diagnostic analysis of images and new diagnostic criteria not currently possible. The methods disclosed in this application provide improved accuracy and greater ease in all of these potential applications.
Radiologists routinely make subjective and even quantitative measurements of foreign masses, tissues or organs by manually placing cursors to define the 2-D extent of the target. If the window and/or level (brightness and contrast) are changed in the display, the apparent size of the target changes because the boundary is not discrete and is moved in or out of the display range. Thus, the measured object size is frequently inaccurate and will vary from operator to operator and from scanner to scanner depending on the display conditions and scanner properties. In addition, the process to set and adjust the window and level requires operator time and is currently very inefficient. Electronic image data are frequently erased, and only the films retained for the medical records.
Prior art methods have allowed CT scanners to be used as quantitative instruments for bone density measurements in quantitative computerized tomography (QCT) by the use of calibration phantoms scanned simultaneously with the patient (simultaneous calibration). Such phantoms have greatly aided the standardization accuracy and reproducibility of bone density measurements. Representative prior art methods have been disclosed in U.S. Pat. No. 4,233,507 to Volz), U.S. Pat. No. 4,985,906 to Arnold, U.S. Pat. No. 5,335,260 to Arnold, U.S. Pat. No. 4,782,502 to Schulz, U.S. Pat. No. 4,651,335 to Kalender, U.S. Pat. No. 4,870,666 to Lonn, and others. These prior phantom designs included, for example, samples within pillows and positioned beside the subject; a flexible phantom positioned on top of the subject; a flexible phantom positioned under the subject with means to force contact to patient; samples within a slot between two couch pads; and in a removable, rigid structure or table top section in the original Volz patent.
The Volz patent used samples with 5 cm cross-sectional diameters and subsequent improvement patents used similar or larger sample areas because of the need to have sufficient numbers of pixels in each CT slice region of interest (ROI) for statistical measurements. The base material and calibration samples were large and dense and increased the dose and image artifacts. U.S. Pat. No. 4,985,906 to Arnold discloses improvements to prior phantoms by using tissue equivalent base material with reduced size and mass to minimize beam hardening effects.
In U.S. Pat. No. 6,990,222 to Arnold discloses a method for hybrid calibration using simultaneous phantom calibration along with internal tissue references of the individual patient. U.S. Pat. No. 6,990,222 is incorporated by reference herein in its entirety. More recently, fast multi-detector MDCT scanners have been used for coronary and aortic calcium analysis with or without simultaneous phantom calibration.
The prior calibration methods were successful because CT numbers, (Hounsfield Units, HU), are only estimates of the x-ray attenuation coefficients of tissue relative to water as the manufacturer's calibration reference material. CT numbers fail to be truly quantitative for several reasons. For example, the tissue attenuation coefficients are photon energy dependent, and the x-ray beam energy spectra are not measured or known for individual patients. Further, many beam energy spectra exist in each CT slice (e.g., a unique spectrum for each path length through the patient and seen at a particular detector element, and a unique spectrum for each view through the patient). The beam spectrum changes with the thickness and composition of tissues in the path length. The quantities of fat, soft tissue, air, and bone vary with each projection. X-ray tube filtration to shape the beam intensity also changes the beam spectrum resulting in variations in tissue attenuation based on locations within the field of view. Scattered radiation is also variable and dependent on some of these parameters. Manufacturers' calibrations (for example, CT number and beam hardening corrections in current practice and scatter correction) are based on idealized phantoms scanned independently from the patient, which are often circular in shape and composed of water, plastics, or other homogeneous, synthetic materials. These differ significantly from the shape and varied composition of real patients. Image pixel intensities vary from image to image, and are dependent on table height, position in the beam, scanner drift, tube changes, manufacture reconstruction software, body region thickness and volume, field of view, and sometimes even the time of day as the imagers warm up.
Measurement errors also result from organ motion artifacts, tissue heterogeneity (fat, muscle, blood mixtures in sub-voxel volumes), gantry wobble and vibrations of the x-ray source, escape of K x-rays in the detectors, reconstruction algorithm, electronic noise, local beam hardening, scattered radiation and structure artifacts (usually from the presence of larger bones, air volumes and table structures). As a result of the measurement errors, tissue densities vary with imaging device, image acquisition parameters, gating and motion, etc., some of which have been reported. In addition, measurements will vary with patient size and body type which is poorly recognized. Because of the finite detector element size, the x-ray tube focal spot size and geometry, source movement during scanning and image back projection reconstruction imperfections, the finite number of views and source and/or table movement, the reconstruction of objects is not exact resulting in edge blurring by the point spread function (PSF) and loss of accurate tissue density representations at edges of objects. As a result, bone density and calcium scores vary significantly with different devices, over time and between different institutions, patient body composition and imaging techniques. The above-referenced measurements are also dependent on this listing of sources of errors.
Prior methods for simultaneous CT calibrations in bone density measurements have used calcium phantoms which are removed after each exam. This requires the operator to place the phantom on the CT table before and to remove the phantom after each quantitative study and then place the phantom in storage. Further, the operator is required to position the phantom in a surrounding foam pad for patient comfort and then to position the phantom and pad and patient in the CT field-of-view (FOV). It is important to reproduce the position of the phantom in the FOV and to avoid its movement during the scans. Misalignments cause variable results and failure to reproduce positioning on follow-up exams months or years later can reduce measurement precision. Unlike these prior methods, the device and methods disclosed in the present application overcomes all of these limitations. Further, the operator does not need to be concerned with phantom positioning or repositioning and thus saves time during each patient procedure.
With prior methods, any study without the phantom cannot then or later be analyzed. With the device and method disclosed herein, any prior study from months to years can be analyzed at any later date since the references may always be present in all images and all studies. Increased radiation doses, costs, times and efforts of repeating scans for quantitative analysis are avoided when existing scans made for other exams can be reprocessed. Since many medical conditions such as osteoporosis, coronary or vascular disease, emphysema, body composition, etc., are chronic and slowly changing conditions, later measurements are likely to remain highly relevant even if analyzed months or years later.
Prior art methods employ large cross-sectional area phantoms with higher density base materials. The large sample sizes were required because software methods operated on usually 2 to 4 reconstructed thick slices by manual or later automated methods. (See, for example, U.S. Pat. No. 4,922,915 to Arnold.) These features are now known to cause beam hardening and scatter artifacts and cause streaks from the reference phantom itself. The larger/dense phantom increases radiation dose to the patient and/or reduces image quality by adding noise and scatter. Because of the automated software methods disclosed herein, much smaller, cross-sectional reference samples can be used with great advantage.
Prior methods and phantoms create patient discomfort from the hard material and sharp edges as well as large size. Patient discomfort often leads to patient motion during the scan creating additional errors. The device and method disclosed herein overcome patient discomfort. Automated software methods to locate and measure reproducibly the very small diameter samples were not available in prior methods. Further, finding and measuring these small samples which are now not rigid and which extend over long distances required advanced methods disclosed herein. With the advent of MDCT scanners with several hundred CT images per exam, calibration on all images in background automatically was required.
Prior methods used phantoms of reduced length, which were significantly shorter than the length of the patient's torso. The required large sample sizes and removal of the phantom between studies resulted in short phantoms. The phantom had to be repositioned between multiple scans, such as hip and thoracic vertebral BMD scans. In addition, the operator must be aware of the position of the phantom and/or scout scan the patient to verify phantom positioning under the anatomical region to be scanned. This requires added time and radiation dose to the patient. The disclosed devices and method overcome these limitations.
Prior methods have been limited primarily to bone density measurements using calcium equivalent phantom samples. Additional reference and calibration materials and new calibrations and image corrections are possible with the currently disclosure method and system.
Prior methods to correct CT scanners for gantry, tube and table motion and methods to measure or correct for image reconstruction, detector variations, system motion and wobble, have used phantoms with pins, spheres and small diameter wires contained within idealized phantoms. These phantoms and test objects have been scanned independent of the patients. These methods, while providing the ability to correct for scanner imperfections in idealized conditions, are not necessarily representative of the imaging device when the patients are present and are not representative of the device at the time of specific clinical exams. X-ray imagers and CT scanners are known to drift, change sensitivity with use and environmental conditions, patient size and weight, (CT table movement, backlash and bending), and scanner and scanning parameters. For all of these and additional reasons, current test phantoms and methods for image corrections have limitations because they have been applied without the patient present. The disclosed device and methods by including similar phantom targets for image corrections simultaneously with each individual patient exam and at the relevant time of the exam overcomes many of these limitations.